Hearing loss, which may be due to many different causes, is generally of two types, conductive and sensorineural. Of these types, conductive hearing loss occurs where the normal mechanical pathways for sound to reach the hair cells in the cochlea are impeded, for example, by damage to the ossicles. Conductive hearing loss may often be helped by use of conventional hearing aid systems, which amplify sound so that acoustic information does reach the cochlea and the hair cells.
In many people who are profoundly deaf, however, the reason for deafness is sensorineural hearing loss. This type of hearing loss is due to the absence of, or destruction of, the hair cells in the cochlea which transduce acoustic signals into nerve impulses. These people are thus unable to derive suitable benefit from conventional hearing aid systems, because there is damage to or absence of the mechanism for nerve impulses to be generated from sound in the normal manner.
It is for this purpose that prosthetic hearing implant systems have been developed. Such systems bypass the hair cells in the cochlea and directly deliver electrical stimulation to the auditory nerve fibres, thereby allowing the brain to perceive a hearing sensation resembling the natural hearing sensation normally delivered to the auditory nerve. U.S. Pat. No. 4,532,930, the contents of which are incorporated herein by reference, provides a description of one type of traditional prosthetic hearing implant system.
Prosthetic hearing implant systems have typically consisted of two key components, namely an external component commonly referred to as a processor unit, and an implanted internal component commonly referred to as a stimulator/receiver unit. Traditionally, both of these components have cooperated together to provide the sound sensation to an implantee.
The external component has traditionally consisted of a microphone for detecting sounds, such as speech and environmental sounds, a speech processor that converts the detected sounds and particularly speech into a coded signal, a power source such as a battery, and an external antenna transmitter coil.
The coded signal output by the speech processor is transmitted transcutaneously to the implanted stimulator/receiver unit situated within a recess of the temporal bone of the implantee. This transcutaneous transmission occurs through use of an inductive coupling provided between the external antenna transmitter coil which is positioned to communicate with an implanted antenna receiver coil provided with the stimulator/receiver unit. This communication serves two essential purposes, firstly to transcutaneously transmit the coded sound signal and secondly to provide power to the implanted stimulator/receiver unit. Conventionally, this link has been in the form of a radio frequency (RF) link, but other such links have been proposed and implemented with varying degrees of success.
The implanted stimulator/receiver unit typically includes the antenna receiver coil that receives the coded signal and power from the external processor component, and a stimulator that processes the coded signal and outputs a stimulation signal to an intracochlea electrode assembly which applies the electrical stimulation directly to the auditory nerve producing a hearing sensation corresponding to the original detected sound.
The external componentry of the prosthetic hearing implant has been traditionally carried on the body of the implantee, such as in a pocket of the implantee's clothing, a belt pouch or in a harness, while the microphone has been mounted on a clip mounted behind the ear or on a clothing lapel of the implantee.
More recently, due in the main to improvements in technology, the physical dimensions of the speech processor have been able to be reduced allowing for the external componentry to be housed in a small unit capable of being worn behind the ear of the implantee. This unit has allowed the microphone, power unit and the speech processor to be housed in a single unit capable of being discretely worn behind the ear, with the external transmitter coil still positioned on the side of the implantee's head to allow for the transmission of the coded sound signal from the speech processor and power to the implanted stimulator unit.
This need for a transmitter coil further requires leads and additional componentry which add to the complexity of such systems as well as being quite noticeable. Nevertheless, the introduction of a combined unit capable of being worn behind the ear has greatly improved the visual and aesthetic aspects for prosthetic hearing implant implantees and provided a degree of freedom of movement for implantees that had previously not been possible with body worn devices.
While traditional prosthetic hearing implants have proven very successful in restoring hearing sensation to many people, the construction of the conventional implant with its external electronic components has limited the circumstances in which the implant can be used by a implantee, For example, implantees cannot wear the devices while showering or engaging in water-related activities. Most implantees also do not use the devices whilst asleep due to discomfort and the likelihood that the alignment between the external transmitter coil and the internal receiver coil will be lost due to movements during sleep. Therefore, with the increasing desire of prosthetic hearing implant implantees to lead a life that is the equivalent of a naturally hearing person, there exists a need to provide a system which allows for total freedom with improved simplicity and reliability.
Because of this need, fully implantable systems that do not require external componentry for operation have been postulated. One type of system which has been proposed is described in U.S. Pat. No. 6,067,474 by Advanced Bionics Corporation and Alfred E Mann Foundation for Scientific Research. This system attempts to provide all system components implanted in the implantee, and includes a microphone placed in the ear canal which communicates with a conventionally positioned stimulator unit via a conventional RF link. There is further described a battery unit which can be integral with the stimulator unit or separate therefrom. Such a system provides further complications as it requires surgical implantation of a number of components and hence complicates the surgical procedure The system also maintains the need for an RF link during normal operation between implanted components which increases overall power requirements of the system and unnecessary drains the internal battery supply. Also, such a system requires multiple implanted casings and the necessity for communications between internal components thereby increasing the likelihood of system failure due to component malfunction. In the event of a system malfunction, the procedure required to correct such a device failure becomes further complicated due to the number of implanted components and the complex communication channels between all components.
The present applicant has also proposed a totally implanted prosthetic hearing implant system in International Application No. PCT/AU01/00769. This system has the advantage that all of the components are provided in a single unit that is able to be implanted by conventional surgical procedures.
A problem with totally implanted system is that the systems are reliant on power sourced from rechargeable power sources implanted with the implant. A well understood problem with rechargeable batteries is that-the batteries can only undergo a particular maximum number of recharging cycles before the performance of the battery degrades to a level where the battery is essentially unusable A further problem with rechargeable batteries is that if they are discharged before being fully charged, or conversely, are charged before being fully discharged, their overall capacity may be reduced. When such batteries are implanted, all means should be undertaken to maximize the batters operating life as the replacement of such batteries requires a surgical procedure.
These problems are further compounded by the fact that in prosthetic hearing implant applications, the recharging process for the implanted power source is heavily reliant on the implantee's ability to dedicate a particular amount of their time to recharge the internal power source when necessary. This requires the implantee to closely monitor the charge status of their system and ensure that the power source is recharged only when the battery has been fully discharged. Therefore, such an onus can impinge greatly on the implantee's lifestyle, thereby reducing the implantee's freedom to use such a totally implanted device, which is one of the great benefits of providing a totally implanted device in the first instance.
It is therefore an object of the present invention to provide a system designed to maximize the performance of batteries installed in totally implanted prosthetic hearing implants and other implants that may rely on battery power.
It is a further object of the present invention to provide an internal power management system that ensures that the charging/recharging cycles of the implanted device occur with minimal interruption to the implantee's regular lifestyle and with minimal input from the implantee.